Method and apparatus for shielding a linear accelerator and a magnetic resonance imaging device from each other

ABSTRACT

A radiation therapy system comprises a magnetic resonance imaging (MRI) system combined with an irradiation system, which can include one or more linear accelerators (linacs) that can emit respective radiation beams suitable for radiation therapy. The MRI system includes a split magnet system, comprising first and second main magnets separated by gap. A gantry is positioned in the gap between the main MRI magnets and supports the linac(s) of the irradiation system. The gantry is rotatable independently of the MRI system and can angularly reposition the linac(s). Shielding can also be provided in the form of magnetic and/or RF shielding. Magnetic shielding can be provided for shielding the linac(s) from the magnetic field generated by the MRI magnets. RF shielding can be provided for shielding the MRI system from RF radiation from the linac.

RELATED APPLICATION

This application is a continuation and claims benefit of priority under35 U.S.C. §120 of U.S. patent application Ser. No. 12/837,309, filedJul. 15, 2010, entitled “Method and apparatus for shielding a linearaccelerator and a magnetic resonance imaging device from each other”,now U.S. Pat. No. 8,836,332 issued Sep. 16, 2014, which claims thebenefit of priority under 35 U.S.C. §119 of U.S. Provisional ApplicationNo. 61/225,771, filed Jul. 15, 2009, titled “Method and apparatus forshielding a linear accelerator and a magnetic resonance imaging devicefrom each other,” which is hereby incorporated by reference.

BACKGROUND

1. Technical Field

The present application relates to systems and methods for combinedradiotherapy and magnetic resonance imaging, particularly systems andmethods that involve shielding magnetic fields and radiofrequencyradiation from the radiotherapy and magnetic resonance imaging systems.

2. Related Art

A linear particle accelerator (also called a linac) is a type ofparticle accelerator used to accelerate subatomic ions at great speeds.Linacs are described, for example, by C. J. KARZMARK ET AL., MEDICALELECTRON ACCELERATORS (McGraw-Hill, Inc., Health Professions Division1993), which is hereby incorporated by reference. Medical grade orclinical linacs (a.k.a. clinacs) accelerate electrons using atuned-cavity waveguide in which the Radio frequency (RF) power typicallycreates a standing or traveling wave for the generation of high energyelectrons or Bremsstrahlung X-rays for medicinal purposes.

Magnetic Resonance Imaging (MRI), or nuclear magnetic resonance imaging(NMRI), is primarily a medical imaging technique most commonly used inradiology to visualize the internal structure and function of the body.MRI is described, for example, by E. MARK HAACKE ET AL., MAGNETICRESONANCE IMAGING: PHYSICAL PRINCIPLES AND SEQUENCE DESIGN (Wiley-Liss1999), which is hereby incorporated herein by reference.

It is desirable to be able to image with an MRI unit while being able tosimultaneously perform radiation therapy using a linac. However, thereare two major conflicts between the linac and MRI unit that should beovercome for these technologies to work together in a clinicallyacceptable way. The first problem is that the MRI unit's magnetic fieldaccelerates charged particles in the linac by the Lorentz force oncharged particles in a magnetic field determined by the equation F=q(v×B), where F is the force on the charged particle, q is the charge ofthe particle, v is the velocity, and B is the magnetic field. In linearaccelerators, the electrons “ions” are typically generated by heating athermionic material (a material where the electrons become detached whenheated), which is the cathode, and when a positive voltage is applied toan anode (which is typically a wire grid), the electrons move from thecathode towards the anode. The anode is pulsed at 100's of megahertzsuch that the grouping of electrons pass thru the grid and on to befurther accelerated. The cathode, anode, and later acceleratingcomponents form what is called the electron gun, and this gun can beshut down by an external magnetic field such that it will not produceelectrons for further acceleration. The MRI magnet is usually shieldedto reduce the magnetic field surrounding the magnet. Usually thismagnetic fringe field remains above the level of the earth's 1 gaussmagnetic field for a few meters from the MRI isocenter. The optimaldistance for locating a linac near the patient is with the source atapproximately one meter from the radiotherapy isocenter. For a systemwhere the MRI and radiotherapy isocenters are substantially coincident,this puts the linac in a fringe field that could easily be on the orderof 0.1 tesla (T, 1T=10,000 gauss) or higher. The magnetic field B vectoris significant and oriented axial to the MR system (Z). The velocity vvector approaches the speed of light and is nominally at right angles(Y) to the B vector. The force F on the very light electron willaccelerate the electrons perpendicularly out of their desiredtrajectory.

The second problem is that the high-powered RF source of the linaccauses interference with the radiofrequency transmitter and receiver forsignal detection in the MRI unit. The RF frequency transmit and(especially) receive coils employed are extremely sensitive and usuallylimited by thermal noise in the patient and RF coil structure. Gradientmagnetic fields are used to set a range of frequencies around thiscentral frequency to provide position information as a function offrequency. The high-powered RF source in the linac typically generatesmegawatt to tens of megawatt bursts of RF radiation tuned to theresonating cavity of the accelerator at several hundred Hertz duringoperation. This high-powered RF radiation is typically not on resonancewith the MRI frequencies of operation, but has side bands at the MRIfrequencies and can induce eddy currents in the conducting components ofthe MRI causing signal corruption or even damaging the MRI electronics.MRI systems usually include an RF shielded room to limit interferencefrom external RF sources. The sensitive MRI receive-RF coils also needto be protected from the RF transmit field used for excitation. Usuallythis isolation is done with PIN diodes and/or back-to-back diodes,switching in/out tuned/detuned circuit elements that attenuate the RFinduced signal. Further, it is important that the sensitive MRI pre-ampsdo not go into saturation with RF energy from any source.

U.S. Pat. No. 6,198,957 to Green, titled “Radiotherapy Machine IncludingMagnetic Resonance Imaging System” (hereinafter “Green”), teaches thecombination of a MRI system and a horizontal linac. Green teaches thatDC coils should extend around the horizontal linac to shield the MRIfrom magnetic fields produced by the linac and that DC coils should beused around the MRI to shield the linac from the leakage magnetic fieldof the MRI. Also, Green teaches that, for a linac that uses an electronbeam, the main magnets of the MRI must be pulsed off while the electronbeam of the linac is pulsed on. In an analogous way, PCT InternationalPublication WO2004/024235 to Lagendijk et al., titled “MRI in GuidedRadiotherapy Apparatus with Beam Heterogeneity Compensators”(hereinafter “Lagendijk”), teaches integrating DC coils into the designof the main magnet of the MRI to create a toroidal low field regionoutside the MRI to shield the linac electron gun source from the MRIleakage magnetic field. Lagendijk also teaches the design of a mainmagnet that provides limited shielding on the electron gun of the linacand allows higher fields along the accelerating path toward the linactarget, though this permits further degradation of the beam and thatrequires correction with additional filters. Again, in a similar way,PCT International Publication WO2007/045076 to Fallone et al., titled“Integrated External Beam Radiotherapy and MRI System” (hereinafter“Fallone”), teaches that a shielding interface between the MRI and linaccan be used if the linac and MRI are fixed to each other to allowshimming, as was disclosed by Green. Also, Fallone teaches the use ofsteering coils associated with the linac for detecting and correctingfor deviations of the linac electron beam due to the magnetic field ofthe MRI. Finally, U.S. Patent Application Publication 2008/0208036 toAmies et al., titled “Combined Radiation Therapy and Magnetic ResonanceUnit” (hereinafter “Amies”), teaches that the linac can be placedcompletely inside the MRI main magnet bore with the path of theaccelerated electrons aligned with the main magnetic field lines,however, this shortens the distance of the linac from isocenter. Thisalso limits the beam path to be exactly along the central axis of themagnet. In a horizontal bore magnet, the magnetic field lines begin todiverge away from the central axis as you approach either end of themagnet, and in so doing turn in a radial direction. Thus, the beam mustbe exactly along the central axis or else it will be effected by theradial components of the field toward the ends. The MRI also uses“pulsed gradient fields” which can also have significant radialcomponents off the central axis. Each of these references also teach theshielding of the linac from the MRI magnetic field where shieldingmaterial is interposed or interfacing between the beam source and thepatient.

Prototypes of the devices taught by Lagendijk (and related itsapplications) and Fallone have demonstrated that the shielding leads tovery large devices that cannot fit in the standard linac (or clinac)room and present many technical challenges where significant compromisesmust be made in the quality of the radiotherapy that can be delivered,either requiring the radiotherapy devices to treat from large distancesor through a large amount of material that can scatter and attenuate thebeam, compromising the quality of the radiotherapy. Additionally, theseprototypes have employed RF shielding boxes that completely enclose theMRI from the linac and the treatment room, making patient access anissue.

As will be appreciated, there exists a need for an improved solution tothe shielding of an MRI and linac from each other that, among otherthings, mitigates the disadvantages of having to pass the radiotherapybeams through a large amount of material or from long distances.

SUMMARY

Disclosed herein are methods and apparatus embodiments that allow forthe production of a combined linac and MRI device. A method of shieldingthe RF radiation of the linac without sealing off the MRI is alsodescribed. Embodiments disclosed herein describe shielding to isolatethe linac from the magnetic field of the MRI magnet and the RFtransmit/receive coils from the linac RF field. A novel method ofshielding the linac from the leakage magnetic field of the MRI at thestandard position, i.e., about one meter from the radiotherapyisocenter, without placing shielding material between the patient andthe incident beam, thereby preventing the degradation of the beam, istaught with shimming and correction of the homogeneous MRI magneticfield with gantry and MRI bore mounted shims; the gantry mounted shimsbeing able to rotate with the linac. Magnetic shielding can be done withferromagnetic shields and local coils, or combinations thereof, that areplaced around the linac, yet not in the path of the beam. RF shieldingof the MRI system is achieved by the selective use of a combination ofuniform RF radiation absorbing materials, such as carbon fiber mesh, andRF radiation reflective materials, such as copper shielding. The beam isallowed to pass through the RF shielding as it can be constructed to bepart of the flattening filter attenuation or can be made with a thinsection or hole to pass the beam. The absorbing and attenuatingmaterials can be layered successively to reflect, attenuate, and/orabsorb the RF radiation from the linac. Cooling can be provided to theabsorbing material as necessary to remove heat generated by the RFradiation being absorbed.

According to some aspects of the present disclosure, a magnetic shieldcan be provided about a linac. The shield can include one or more shellsof high magnetic susceptibility and permeability layers, currentcarrying coils, permanent magnets, or any combination thereof, to shieldthe linac from the magnetic field of a MRI system in order to allow forproper operation of the linac. The shells are preferrably cylindrical,but other shapes can be used.

In embodiments that include more than one of the shells, the shells arepreferrably magnetically isolated from each other.

The shield can be arranged so that the magnetic field of the MRI systemdoes not attenuate the radiotherapy beam. The shield can operate at apreferred distance for linac placement. The inner layers of the shieldcan have higher permeability but saturate at a lower flux density. Theinfluence of the shield on the homogeneous region of the MRI magneticfield can be diminished and balanced by an opposed dummy shield.

The influence of the shield on the MRI magnetic field can be correctedby shims. For example, gantry mounted shims can correct perturbationsthat follow the gantry angle of linac. MRI bore mounted shims and/ormagnet design can correct for perturbations that are independent of thegantry angle of the linac.

According to further aspects of the present disclosure, an RF shieldabout a linac can include one or more layers of RF absorbing materials,and/or RF reflecting materials, or combinations of both, to contain theRF radiation and/or shield the MRI from the high power RF radiationproduced by the linac in order to allow for proper operation of the MRI.

The RF shield can be arranged so that the beam passes through the shieldwith uniform attenuation. The RF shield can also be arranged so that theflattening filter is part of the RF shield. A thin section or hole canbe used to limit beam attenuation.

Shielding can be improved by the application of RF absorbing materialsto one or more of the RF room interior walls, the MRI surfaces, and theformer for winding the gradient coils.

BRIEF DESCRIPTION OF THE DRAWINGS

Features, aspects, and embodiments of the inventions are described inconjunction with the attached drawings, in which:

FIG. 1A shows a plan view of a split-magnet radiation therapy system;

FIG. 1B shows a perspective view of the split-magnet radiation therapysystem shown in FIG. 1A;

FIG. 1C shows a simplified block diagram of the split-magnet radiationtherapy system shown in FIG. 1A;

FIG. 1D shows another simplified block diagram of the split-magnetradiation therapy system shown in FIG. 1A;

FIG. 1E shows another simplified block diagram of the split-magnetradiation therapy system shown in FIG. 1A;

FIGS. 2A and 2B shows charts of a magnetic field generated by mainmagnets of an MRI of the system shown in FIGS. 1A-1D;

FIGS. 3A and 3B show charts of the B-H curve and the relativepermeability, respectively, of magnetic shielding material used in someembodiments of the system shown in FIGS. 1A-1D;

FIG. 4A shows a simplified block diagram of some embodiments of thesystem shown in FIGS. 1A-1D, including a section view of the mainmagnets shown in FIGS. 1C and 1D;

FIGS. 4B-4E show more detailed views of embodiments of the shieldingthat can be used with the system shown in FIGS. 1A-1D;

FIGS. 5A and 5B show a comparison of the shielded and unshieldedBz-field generated by the main magnets of the MRI according to someembodiments;

FIGS. 6A and 6B show a comparison of the shielded and unshieldedBz-field generated by the main magnets of the MRI according to thepreferred embodiment;

FIGS. 7A and 7B show the Bz-field map inside the preferred embodiment ofa magnetic shield in the XY planes at Z=10 mm and Z=20 mm respectively;

FIG. 8 shows a simplified block diagram of some embodiments of thesystem shown in FIGS. 1A-1D, including a section view of the mainmagnets shown in FIGS. 1C and 1D;

FIG. 9 shows the Bz-field generated by the main MRI magnets as shieldedaccording to some embodiments;

FIG. 10 shows a simplified block diagram of some embodiments of thesystem shown in FIGS. 1A-1D;

FIG. 11 shows an embodiment of an active shield that can be used in someembodiments of the system shown in FIGS. 1A-1D; and

FIGS. 12A-12B show the z-component of the magnetic field generated bythe main MRI magnets before and after, respectively, activation of theactive coil shown in FIG. 11.

DETAILED DESCRIPTION

FIGS. 1A-1E show various views of a split-magnet radiation therapysystem 100. FIGS. 1A and 1B show plan and perspective views,respectively, of a split-magnet radiation therapy system 100. The system100 includes an integrated linear accelerator 107 and MRI system 102,and allows for simultaneous irradiation from the linear accelerator 107and imaging from the MRI 102. For example, the MRI 102 can be used topinpoint the location of an object to be irradiated, and thisinformation can be used to control the irradiation from the linearaccelerator 107. The present disclosure is not necessarily limited tothe specific MRI and linac systems shown in the Figures and referencedherein, but can apply equally to other MRI and linac systems. Forexample, RF and/or magnetic shielding systems and methods disclosedherein can be used with known MRI and linac systems that may differ fromthose shown in the Figures and described below.

The radiation therapy system 100 includes an open split solenoidalmagnetic resonance imaging (MRI) device 102, a radiation source 104, agantry 106 for housing a linac 107 and for changing the angle ofradiation source 104, a patient couch 108, and a patient 110 in positionfor imaging and treatment. A similar system is described in U.S. PatentApplication Publication 2005/0197564 to Dempsey, titled “System forDelivering Conformal Radiation Therapy while Simultaneously Imaging SoftTissue” (hereinafter “Dempsey '564”), which is hereby incorporated byreference.

The radiation therapy system 100 of the present disclosure differs inmany respects from that disclosed in Dempsey '564, a primary differencebeing that the radiation therapy system 100 of the present disclosureincludes a linac 107 rather than the isotopic radiation system disclosedin Dempsey '564. Except as described herein, the linac 107 can be ofconventional design. In some embodiments, the linac 107, best shown inFIG. 1E, can be a medical grade or clinical linac (clinac) configured toaccelerate electrons using a tuned-cavity waveguide 107 a in which theRadio frequency (RF) power creates a standing or traveling wave for thegeneration of high energy electrons from an electron gun 107 b. Anoptional target 107 c can be included that is installed forx-ray/photon-beam therapy and removed for electron-beam therapy. TheX-ray/photon beams and electron beams constitute examples of linacradiation beams. In some embodiments, the system 100 can include apre-collimator 107 d and a multi-leaf collimator 107 e, for example asdisclosed in Dempsey '564, for the electron beam EB from the linac 107.As discussed in greater detail below, the linac 107, particularly thewaveguide 107 a, can be protected by magnetic and/or RF shielding 118,120, and/or 122. The magnetic and/or RF shielding 118, 120, and/or 122can be in the form of one or more shells that are preferrablycylindrical, but other shapes can be used. Also, as discussed in greaterdetail below, the radiation therapy system 100 can include a coolingsystem 115 for cooling the shielding 118, 120, and/or 122. The coolingsystem 115 can include, for example, liquid and/or air cooling systems.

The radiation therapy system 100 can include a split magnet system, suchas described in Dempsey '564. The split magnet system includes a pair ofmain magnets 112 a and 112 b as shown in FIG. 1C as part of the MRIdevice 102. The MRI device 102 can also include conventional MRIcomponents that are not shown, such as a split gradient coil, one ormore shim coils (also referred to as shims), and an RF system, includingRF coils. The strength of the magnetic field generated by the mainmagnets 112 a and 112 b can vary. However, for convenience ofexplanation, the system 100 will be described with reference to anembodiment where the main magnet field strength is 0.35 T, which ischosen to prevent perturbations in the dose distribution caused by theLorentz force acting on secondary electrons in the patient. The magnets112 a and 112 b are separated by a central gap 114, for example of 0.28m. The MRI device 102 can be designed to provide an MRI field-of-viewof, for example, 50 cm diameter around a center of the image field, andat the same time provide an un-attenuated radiation beam in the gap 114with the split gradient coil of the MRI device 102. Preferrably, thesystem 100 is constructed such that the radiation beam from the splitgradient coil only passes through RF coils, the patient 110, and thepatient couch 108.

FIGS. 1C and 1D show a simplified block diagrams of the system 100. InFIG. 1C, only the main magnets 112 a and 112 b of the MRI system 102 areillustrated; in FIG. 1D only the main magnets 112 a and 112 b and thelinac 107 are illustrated. The coordinate system shown in FIGS. 1C and1D, and used throughout this disclosure, refers to the longitudinal axisthrough the MRI bore (lengthwise through patient 110) as the Z-axis. TheZ-axis is normal to a central axial plane CP, also referred to astransverse or central plane CP, which is at least substantially centeredwithin the gap 114 between the main magnets 112 a and 112 b. Also, themain magnets 112 a and 112 b both extend radially about the Z-axis. Thecentral plane CP is also defined by an X-axis and a Y-axis. The X-axisextends perpendicular to the Z-axis and from side to side of the MRIsystem 102; the Y-axis extends perpendicular to the Z-axis and frombottom to top of the MRI system 102.

In the system 100 of the present embodiment, at a distance of 1 m frommagnet isocenter IC on the central plane CP, there is a magnetic fieldof Bz≈0.1 T, shown as point P1, which is a desired distance fromisocenter for the source of the radiation of the linac 107. The magneticfield reverses direction from +Bz to −Bz at a radial distance of 0.81 m,shown as point P2. The magnet field at 1 m from isocenter, where thelinac 107 radiation source is preferrably located for optimalradiotherapy operation, is low enough that it can be contained in aferromagnetic shield or multiple layered shields, as described below. Inthe central axial plane CP, there is mainly axial magnetic field Bzbecause of coil symmetry. In the central plane CP, we assume that Y is avertical axis and the axis of a high magnetic susceptibility (and/orpermeability in a linear domain) material, e.g., a non-orientedsilicon-steel shell, for shielding the linac 107.

The field generated by the main magnets 112 a and 112 b near the centralplane CP is shown in FIG. 2A. For the linac 107 to operate properly, itis desirable for the magnetic field in the center of the acceleratingstructure to be much less than the unshielded magnetic field near Y=1000mm (e.g., point P1). FIG. 2A also shows that there is a null point(Bz=0) in the vicinity of yz≈810 mm where the Bz field reversesdirection, as must always happen due to the reversal of field direction.FIG. 2B shows the same field in the region of interest near Y=900 mm,but with a rescaled Y-axis.

The linac 107 has a longitudinal axis p that is aligned with the Y-axisin FIG. 1D. While the linac 107 is shown and described as being alignedalong the Y-axis, it is preferable for the linac 107 to be rotatableabout the Z-axis. For example, the gantry 106 shown in FIGS. 1A and 1Bcan support the linac 107 and carry the linac 107 about the Z-axis(while the longitudinal axis p remains in the central plane CP), in therotation directions RD shown in FIG. 1D, such that the linac 107 canemit an electron beam EB towards the isocenter IC from any, or a rangeof, rotational positions about the Z-axis. Also, the gantry 106 andlinac 107 can rotate about the Z-axis independently of other componentsof the system 100. For example, the gantry 106 and linac 107 can rotateindependently of the MRI 102.

Turning next to FIGS. 3A-3B and FIGS. 4A-4B, we now describe a generalmethod for magnetically shielding the linac 107 from the magnetic fieldof the MRI system 102. Although specific examples are provided, thisdoes not exclude similar approaches or variations in form or material toachieve the same goal. To suppress the magnetic field B in the regionwhere the linac 107 is located, a magnetic shield or shell 118 made ofhigh magnetic susceptibility and permeability material, is placed aroundthe linac accelerating structure 107. The shell 118 can be cylindricalin shape and aligned along axis p of the linac 107, with one or bothends of the shell 118 being open. While a cylindrical shape ispreferred, the disclosed shield shells can be other shapes. At least oneend of the shell 118 is open for the electron beam EB (shown in FIG. 1D)from the linac 107. The magnetic shield 118 can have a thickness chosenaccording to characteristics of the shell material. The magnetic shield118 (as well as other magnetic shields disclosed herein) can be formedof non-oriented silicon steel, for example a nickel-iron alloy, such ascommercially-available material sold by ThyssenKrupp Steel under thetrade name 530-50 AP and having a thickness of, for example, about be 5mm. The B-H curve and relative permeability of “530-50AP” material areshown in FIGS. 3A and 3B, respectively. Other material options for themagnetic shield 118 (as well as other magnetic shields disclosed herein)include M19 steel, M45 steel, and Carpenter High Permeability “49”Steel.

The magnets 112 a and 112 b, and the location of the magnetic shield118, are illustrated in FIG. 4A, while a close-up perspective view ofthe magnetic shield 118 and linac 107 are shown in FIG. 4B. The outerdiameter OD and length L of the magnetic shield 118 can vary; in thepresent embodiment, the outer diameter OD is about 30 cm, and the lengthL is about 70 cm. A bottom edge 118A of the magnetic shield 118 islocated at a fixed distance from the isocenter IC (in the presentembodiment, about 80 cm) that is at or near the Bz field reversallocation, although this is not a requirement. The location and size ofthe magnetic shield 118 should be large enough to contain the linac 107,but not so long or narrow that it limits the size of the beam emitted bythe linac 107. The magnetic shield 118 configuration is optimal forradiotherapy applications when combined with split main magnets 112 aand 112 b and gradient coil set, as the magnetic shield 118 is notimposed between the radiation source of the linac 107 and the patient110. This allows for producing radiotherapy beams of the linac 107 ofhigh quality and strength. In some embodiments, such as shown in FIG.4C, the magnetic shielding can be provided by multiple shield shells. InFIG. 4C, the magnetic shielding is provided by the magnetic shield 118and a second magnetic shield 120, where the shields 118 and 120 can beconcentric layers of steel, which can be separated by layers of air orother insulating material.

The model of the influence of the material, which in this embodiment issteel, in the presence of the main magnets 112 a and 112 b was solvedusing Maxwell's equations via the boundary element method. FIG. 5A showsa comparison of the Bz-field generated by the main magnets 112 a and 112b, and the z-component of the Bz-field generated by the main magnets 112a and 112 b as shielded by an embodiment where the magnetic shieldingcomprises an outer magnetic shield 118 and an inner magnetic shield 120,where the shields 118 and 120 are separated by a layer of air. FIG. 5Bshows a close-up view of FIG. 5A of the region of interest near Y=1200mm of the z-component of the Bz-field generated by the main magnets 112a and 112 b as shielded by the magnetic shields 118 and 120. Table 1lists the materials and dimensions of the magnetic shields 118 and 120according to the embodiment associated with FIGS. 5A and 5B. In Table 1,“ID” is the inner diameter, “OD” is the outer diameter, Length is theshell length L, and the “Starting Y-position” is the distance from theisocenter (Z-axis) to the respective bottom edges of the shields 118 and120.

TABLE 1 Two Shells of 530-50AP Steel Layer ID [mm] OD [mm] Length [mm]Starting Y-position [mm] Inner 260.0 270.0 700.0 900.0 Outer 280.0 300.0700.0 900.0

The residual magnetic field along the axis of a single 5 mm thick shellis about 4.5 G, approximately ten times greater that the earth'smagnetic field and larger than optimal for the linac 107. There areseveral options to further reduce this residual field. As shown in FIG.4C, one option is to add a secondary shielding element 120 inside of themagnetic shield 118 to further reduce the magnetic field that ismagnetically isolated from the first. For example, the secondaryshielding element 120 can be a second shell 120 positioned inside of thefirst shell 118, where both shells are coaxial along the longitudinalaxis ρ of the linac 107. In such embodiments, the second shell 120 canbe of higher permeability, but of a lower saturation flux density of theouter shell 118, as the outer shell 118 has greatly reduced the magneticfield, e.g., mu-metal. It is preferable to magnetically isolate theshells 118 and 120 in order to gain the highest shielding by restartingthe saturation of the metal.

Alternatively, the secondary shielding element 120 can be a currentcarrying coil that is located inside of the primary shell 118 to cancelthe residual field. If the magnetic field remaining is sufficiently lowand its value and direction in space are known, then it can be possibleto make small adjustments in the accelerating portion of the linac Thecurrent linacs are configured to accommodate an electron beam that is atleast substantially straight; if the beam were bent only a small amountby the field, the anticipated beam path can be calculated and theaccelerating plates can be altered to accommodate the beam bending.Given the azimuthally symmetric nature of the fringe field, the pathdeviation of the electron beam should be largely independent of gantryposition. As another alternative, the secondary shielding element 120can be an RF shield 120, as further described below.

The peak-to-peak field in-homogeneity of the system main magnets 112 aand 112 b plus the double shell is 623.8 ppm over 45 cm DSV. Thisinhomogeneity is too large for MRI system 102, so additional shimming isdesirable. The field inhomogeneity is mostly represented by a few of thetesseral harmonics; S_(1,1)→Y, C_(2,2)→(X2-Y2), and S_(3,1)→Z2X, andS_(3,3)→X3. All of the major harmonics of significance are listed inTable 2.

TABLE 2 Spherical Harmonics over 45 cm DSV Zonal Harmonics [ppm]Tesseral harmonics [ppm] n C_(n) n m C_(n,m) S_(n,m) 1 1.625035E−03 1 16.6950990E−03 −2.6417408E+02 2 −9.190121E+01 2 1 −4.3762731E−03−2.2226838E−03 3 4.274773E−03 2 2 −2.3791910E+01 −1.1871930E−03 48.878808E+00 3 1 −1.1657569E−04 1.5830479E+01 5 −2.132553E−03 3 2−1.9884826E−04 5.8882723E−04 6 −6.259163E−01 3 3 −1.0577878E−041.2089904E+00 7 −7.645843E−03 4 1 3.2428894E−04 −2.8578203E−05 83.513474E−01 4 2 8.1373300E−01 3.6183409E−05 9 −9.504502E−03 4 37.2001599E−05 3.3853550E−05 10 2.238179E+00 4 4 4.2607165E−02−5.3185952E−06 11 6.139678E−03 5 1 −2.7178914E−04 −9.0437945E−01

The zonal harmonics can all be handled by shimming, and the shim settingdoes not change with rotation of the linac 107 around the Z-axis. Hence,the shims can be located on the MRI bore. The negative of the zonalharmonics could even be built into the magnets 112 a and 112 b so thatthe combination of magnets 112 a, 112 b plus magnetic shield 118eliminates these terms. The tesseral harmonics are a larger problembecause they would move with the linac orientation. The tesseralharmonics could be shimmed out with passive shims near the central planeCP on the gantry 106 that would move with the gantry 106/linac 107rotation and/or with resistive shims built into the gradient coil thatcould be electrically adjusted to match the rotation of the gantry 106.

According to some embodiments, the system 100 as shown in FIGS. 1A-1Dincludes a linac 107 having a vertical acceleration axis and is mountedon the gantry 106 so that the linac 107 can be rotated about theradiotherapy and MRI 102 isocenters. The linac 107 is also preferred tobe of low energy, in the range of 4 to 6 MV, and have a standing waveguide to keep it compact. The linac 107 can be configured to onlyproduce photon beams that can be used for intensity modulated radiationtherapy or conformal radiation therapy. The linac 107 can operate ateither S-band or X-Band frequencies, but S-band is preferred for highoutput and stability. Referring to FIG. 4C, in this embodiment theelement 120 can be configured to serve as an RF shield 120. In order toprovide RF shielding, the RF shield shell 120 can be made of a suitableshielding material, for example copper foil, aluminum foil, or carbonfiber. Metals such as copper and aluminum tend to reflect RF radiationdue to eddy currents on their surfaces. The carbon fiber materials tendto absorb RF energy.

In some embodiments, particularly where the RF shield shell is formed ofconductive material, the eddy currents can be reduced by providing oneor more slots that extend through the shield shell. For example, shieldshell 120 is shown as having slots 120A and 120B in FIG. 4C. However,the size, number, and configuration of the slots can vary from thatshown in FIG. 4C. Also, while shield shell 120 is shown with slots, suchslots can also, or alternatively, be provided in shield shell 118; also,any number of such slots can be provided in any one or more of theshield shells in embodiments having more than one shield shell. Suchslots can also be desirable in the magnetic shielding shells, and canthus be included in some embodiments of the magnetic shielding shells.

While FIG. 4C shows two layers (shield shells 120 and 118), alternativeembodiments can include any number of layers. In some embodiments, thelayers of shield shells can be made of combinations of differentmaterials or of the same material. For example, in some embodiments, theshield shell layers can include alternating layers formed of RFabsorbing material and RF reflecting material. In such embodiments, itis desirable to provide an air gap between the layers of shield shells.

Cooling can be provided by cooling system 115 (FIG. 1E) as needed to theabsorbing material in the RF shield 120. A variety of known coolingmethods can be used for cooling the RF shield 120. The cooling system115 can include, for example, fluid-carrying conduit for circulating afluid in the vicinity of one or more of the shield shells that form theRF shield 120. Also, air-cooling can be provided by incorporating asystem for moving air across one or more surfaces of the shield shellsthat form the RF shield 120.

The magnetic shield 118 and the RF shield 120 are placed around thelinac 107 to shield the path of the electrons from the electron gun 107b of the linac 107 to the target to a magnetic field strength on theorder of the size of the earth's magnetic field strength. The magneticshield 118 is arranged such that it is not in the path of theradiotherapy beam, for example as shown in FIGS. 4A and 4C. The RFshield 120 is also placed around the linac 107, rather than the MRI 102,and comprised of both absorptive and reflective layers to dissipate andabsorb the RF radiation generated by the linac 107 before it cancompromise the MRI function and they can function as part of theflattening filter. In some embodiments, the RF shield 120 can work inconcert with a standard bore-mounted MRI RF shield. The beam from thelinac 107 is allowed to pass through the RF shield 120 (as well as thebore mounted MRI RF shield in such embodiments) as long as the RFshield(s) are uniformly and minimally attenuating to the radiotherapybeam. It should be noted that in some embodiments, the RF shield 120 canbe provided without the magnetic shield 118 where only the RF shieldingmay be desired.

As mentioned above, in some embodiments, the secondary shielding element120 shown in FIG. 4C can be a second magnetic shield 120. Referring toFIG. 4D, to suppress even further the magnetic B-field in the regionwhere the linac 107 is located, a magnetic shield device 122 can includeone or more concentric magnetic shields, which can include magneticshields 118 and 120 as well as one or more additional magnetic shields.The magnetic shield device 122 can include the multiple magneticshields, including shields 118 and 120, that are made of high magneticsusceptibility (and permeability) material. The shields of the magneticshield device 122 can be concentrically placed inside of each otheraround the linac 107 accelerating structure. The magnetic shields of themagnetic shield device 122 can be magnetically and electrically isolatedfrom each other with a suitable dielectric material such as air orplastic. Having multiple magnetic shields is beneficial because themagnetic field shielding of the material begins to saturate with depth.Introducing a new magnetic shield restarts the saturation effectproviding increased shielding. Also, some embodiments such as the oneshown in FIG. 4E can include a linac 107 having a split radiotherapymagnets 126 and 128 and a magnetic shield made of two isolated shells130 and 132. The thickness of the magnetic shields of the embodimentsshown in FIGS. 4A-4E can be chosen to be, for example, 5 mm, and thematerial can be selected to be 530-50AP steel material. Other materialoptions for the magnetic shield 118 (as well as other magnetic shieldsdisclosed herein) include M19 steel, M45 steel, and steel sold byThyssenKrupp Steel under the trade name 530-50 AP. The outer diameter ODand length L of the shielding shells can be, for example, 27 cm and 30cm, respectively, in a two-shell embodiment such as the one shown inFIG. 4C. The shells 118 and 120 can both be located at a fixed distancefrom the isocenter IC (in the present embodiment, about 85 cm) that isat or near the Bz field reversal location, although this is not arequirement. The location and size of the magnetic shields, includingshields 118, 120, 130, 132, and any additional magnetic shields of themagnetic shield device 122, should be large enough to contain the linac107, but not so long or narrow that it limits the size of the beam fromthe linac 107.

FIG. 6A shows a comparison of the Bz-field generated by the main magnets112 a and 112 b, and the z-component of the Bz-field generated by themagnets 112 a and 112 b as shielded using a magnetic shield device 122that includes three concentric shield shells. FIG. 6B shows a close-upview of the region of interest near Y=1000 mm of the z-component of theBz-field generated by the main magnets 112 a and 112 b as shielded bythe magnetic shield device 122. Table 3 lists the materials anddimensions of the magnetic shield device 122 according to theembodiments associated with FIGS. 6A and 6B. In the embodimentassociated with FIGS. 6A and 6B and Table 3, the magnetic shield device122 includes three concentric shells separated from each other by layersof air. As with other shielding shells disclosed herein, the shells ofthe shield device are preferrably cylindrical, but can be other shapes.In Table 3, “ID” is the inner diameter, “OD” is the outer diameter,Length is the shell length L, and the “Starting Y-position” is thedistance from the isocenter (Z-axis) to the respective bottom edges ofthe layers of the shield device 122.

TABLE 3 Steel M19 and Two Shells of 530-50AP Steel Starting Y- OD Lengthposition Layer Material ID [mm] [mm] [mm] [mm] Inner “M19” Steel 244.0254.0 700.0 900.0 Middle “530-50AP” Steel 260.0 270.0 700.0 900.0 Outer“530-50AP” Steel 280.0 300.0 700.0 900.0

The residual B-field is less than 1 Gauss in the region 1100 mm<y<1400mm. This is roughly comparable to the earth's field close to the axis p.The harmonics of the magnetic field are close to the single shell modelassociated with the embodiment shown in FIG. 4B. The Peak-to-Peak fieldin-homogeneity over 45 cm DSV generated by the main magnets 112 a and112 b plus the magnetic shields 118 and 120 is 623.6 ppm. It ispreferable to have the best shielding on the electron gun 107 b of thelinac 107 and less shielding can be applied to the target end of theaccelerating structure. This field in-homogeneity is mostly representedby the y-harmonic. The spherical harmonics are listed in Table 4.

TABLE 4 Two shells solution: Spherical Harmonics over 45 cm DSV ZonalHarmonics [ppm] Tesseral harmonics [ppm] n C_(n) n m C_(n,m) S_(n,m) 11.6250352E−03 1 1 6.6950990E−03 −2.6417408E+02 2 −9.1901212E+01 2 1−4.3762731E−03 −2.2226838E−03 3 4.2747730E−03 2 2 −2.3791910E+01−1.1871930E−03 4 8.8788081E+00 3 1 −1.1657569E−04 1.5830479E+01 5−2.1325528E−03 3 2 −1.9884826E−04 5.8882723E−04 6 −6.2591632E−01 3 3−1.0577878E−04 1.2089904E+00 7 −7.6458435E−03 4 1 3.2428894E−04−2.8578203E−05 8 3.5134737E−01 4 2 8.1373300E−01 3.6183409E−05 9−9.5045015E−03 4 3 7.2001599E−05 3.3853550E−05 10 2.2381795E+00 4 44.2607165E−02 −5.3185952E−06 11 6.1396783E−03 5 1 −2.7178914E−04−9.0437945E−01

The methods to be used to shim out this inhomogeneity are the same asthose proposed in the case of the single shell model. FIGS. 7A and 7Bshow the Bz-field map inside the inner shell in the XY planes at Z=10 mmand Z=20 mm, respectively.

Referring next to FIG. 8, another embodiment will be described that canreduce field in-homogeneity caused by the presence of a linac shield,such as the shield 118 shown in FIGS. 4A and 4B. The embodiment shown inFIG. 8 can be similar to the embodiment shown in FIGS. 4A and 4B, andlike components have retained the same element numbers; description ofthose components applies equally here, so the description is notrepeated. In the embodiment shown in FIG. 8, the first shield 118extends along a first longitudinal axis ρ1 and a second shield 140(which can optionally include a second linac 107′) extends along asecond longitudinal axis ρ₂ symmetrically 180° apart from the firstlongitudinal axis ρ₁ of the first magnetic shield 118. Each of the axesρ₁ and ρ₂ is on the central plane CP. In some embodiments, the secondshield 140 can be formed of a magnetically shielding material, such assteel sold by ThyssenKrupp Steel under the trade name 530-50 AP, asdescribed in connection with magnetic shield 118. Other material optionsfor the magnetic shield 118 (as well as other magnetic shields disclosedherein) include M19 steel, M45 steel, and Carpenter 49 steel. If only asecond symmetric shield 140 is present, this solution can be thought ofas a symmetric shim for the primary shell 118. In some embodiments, oneor both of the magnetic shields 118 and 140 can be magnetic shielddevices that include two or more concentric magnetic shield shells, suchas shown in FIG. 4C or FIG. 4D.

FIG. 9 shows the Bz-field generated by the main magnets 112 a and 112 band in an embodiment where both the magnetic shield 118 and the magneticshield 140 include two concentric magnetic shielding shells. In thisembodiment, the peak-to-peak field in-homogeneity over 45 cm DSVgenerated by the system main magnets 112 a and 112 b plus the twodouble-shell shield (118+140) is 416.96 ppm. This field in-homogeneityis mostly generated by the Z2 harmonic. The Y-harmonics all becomenegligible small because of the Y symmetry. The harmonics for this caseare listed in Table 4.

TABLE 4 Two double shells solution: Spherical Harmonics over 45 cm DSVZonal Harmonics [ppm] Tesseral Harmonics [ppm] n C_(n) n m C_(n,m)S_(n,m) 1 −1.1158532E−03 1 1 −1.3130497E−04 −1.3130497E−04 2−1.7798728E+02 2 1 9.4937074E−05 9.4937074E−05 3 7.9200018E−03 2 2−4.7129252E+01 −9.2290614E−03 4 1.7600141E+01 3 1 4.5203733E−064.5203734E−06 5 −2.2793685E−03 3 2 −4.0735120E−05 −8.2531950E−04 6−1.3166284E+00 3 3 1.0363288E−05 −1.0363288E−05 7 −1.3414318E−02 4 1−7.1884515E−05 −7.1884515E−05 8 4.0916507E−01 4 2 1.6230890E+002.4395720E−04 9 −1.8969599E−02 4 3 −5.7802678E−06 5.7802678E−06 102.2510390E+00 4 4 8.3827275E−02 1.3021016E−05 11 1.0428939E−02 5 15.3620187E−05 5.3620187E−05

The zonal harmonics are now twice as large as in the single shell modelassociated with the embodiment shown in FIG. 4B. However, they can allbe handled by passive shimming, and the shim setting does not changewith rotation of the linac 107 around the Z-axis. The negative of thezonal harmonics could even be built into the main magnets 112 a and 112b so that the combination of main magnets 112 a and 112 b plus shieldshells 118 and 140 eliminates these terms. The Tesseral harmonics are alarger problem because they would move with the linac 107 rotationalposition. However, symmetry eliminates the worst of the harmonics. TheTesseral harmonics can be shimmed out with passive shims near thecentral plane on the linac gantry 106 and/or with resistive electricalshims. Passive shims built into the rotating gantry 106 can be permanentmagnet shims at these magnetic field levels (oriented magnetizationshims for more shim options). Passive shims can be added at a smallerradius to reduce the material required in the shims. Resistiveelectrical shims in the gradient would change with the rotation of thelinac gantry.

In still further embodiments, there can be N sets of magnetic shieldshells identical or similar to magnetic shield 118, each having arespective axis ρ₁ through ρ_(N). Such embodiments can be arranged in amanner similar to the embodiment shown in FIG. 8. Each of the axes ρ₁through ρ_(N) is on the central plane CP and angularly separated by anangle=360°/N. The higher the N, the more that the net effect of theTesseral harmonics can be canceled out. Also, since the magnetic shieldshells tend to act as RF shields, multiple shells are advantageous forproviding RF shielding.

In some embodiments, as shown in FIG. 10, there can be two parallelannulus discs 144 and 146 made of high relative permeability material.They can be a part of the gantry 106 and on opposing sides of the linac107. In this case, the Tesseral spherical harmonics should be relativelysmall, and the Zonal harmonics should be relatively big. Placing twoannulus discs 144 and 146 in some sense are equivalent to two extracoils in the main magnet 112 a, 112 b. Optimally, the main magnet 112 a,112 b can be designed to accommodate two annulus discs 144 and 146.

The magnetic field from the main magnets 112 a and 112 b at 1 meter fromisocenter along the Y-axis is difficult to shield without the fieldreduction of passive shields, such as shield 118 described above.However, after the magnetic shielding provided by the ferromagneticmaterial, the residual field is near 5-7 Gauss. This residual field caneasily be shimmed out with DC current in a coil, for example inembodiments where the secondary shielding element 120 shown in FIG. 4Cis a coil 120′. A schematic view of the shielding coil 120′ is shown inFIG. 11. The coil 120′ can be cylindrical, having a half-length L andradius R and designed according to the following method (although shapesother than cylindrical can be used). The shielding coil 120′ shouldpreferrably produce the magnetic flux field Bx (in local system ofcoordinates) that cancels the Bz component of the magnetic field (in theoriginal system of coordinates) generated by the main magnets 112 a and112 b.

The current density on the cylinder of radius R can be presented in thefollowing form:{right arrow over (J)}(ρ,φ,z)=δ(ρ−R){ê _(φ) f _(φ)(z)cos(φ)+ê _(Z) f_(Z)(z)sin(φ)}∇{right arrow over (J)}=0

f _(φ)(z)=f _(Z)′(z)

The magnetic potentials generated by this current can be expressed asfollows:

${A_{\rho}\left( {\rho,\varphi,z} \right)} = {\frac{\mu\; R^{2}}{2\;\pi}{\sin(\varphi)}{\int_{0}^{\infty}{k\ {\mathbb{d}{k\left( {{T_{2}\left( {k,\rho,R} \right)} - {T_{0}\left( {k,\rho,R} \right)}} \right)}}{F_{S}\left( {k,z} \right)}}}}$${A_{\varphi}\left( {\rho,\varphi,z} \right)} = {{- \frac{\mu\; R^{2}}{2\;\pi}}{\cos(\varphi)}{\int_{0}^{\infty}{k\ {\mathbb{d}{k\left( {{T_{2}\left( {k,\rho,R} \right)} + {T_{0}\left( {k,\rho,R} \right)}} \right)}}{F_{S}\left( {k,z} \right)}}}}$$\mspace{79mu}{{A_{Z}\left( {\rho,\varphi,z} \right)} = {\frac{\mu\; R}{\pi}{\sin(\varphi)}{\int_{0}^{\infty}\ {{\mathbb{d}{{kT}_{1}\left( {k,\rho,R} \right)}}{F_{C}\left( {k,z} \right)}}}}}$     F_(C)(k, z) = ∫_(L₁)^(L₂)f_(Z)(z^(′))cos (k(z − z^(′))) 𝕕z^(′)     F_(S)(k, z) = ∫_(L₁)^(L₂)f_(Z)(z^(′))sin (k(z − z^(′))) 𝕕z^(′)     T_(n)(k, ρ, R) = θ(ρ − R)I_(n)(kR)K_(n)(k ρ) + θ(R − ρ)K_(n)(kR)I_(n)(k ρ),  n = 0, 1, 2

In this equation, I_(n)(kρ), K_(n) (kρ) are modified Bessel functions.The transverse components of the magnetic field can be presented in thefollowing form:

$\begin{matrix}{{{B_{\rho}\left( {\rho,\varphi,z} \right)} = {{- \frac{\mu\; R^{2}}{\pi}}{\cos(\varphi)}{\int_{0}^{\infty}{k^{2}{\mathbb{d}{k\left( {{{\theta\left( {\rho - R} \right)}{I_{n}^{\prime}({kR})}{K_{n}^{\prime}\left( {k\;\rho} \right)}} + {{\theta\left( {R - \rho} \right)}{K_{n}^{\prime}({kR})}{I_{n}^{\prime}\left( {k\;\rho} \right)}}} \right)}}{F_{C}\left( {k,z} \right)}}}}}{{B_{\varphi}\left( {\rho,\varphi,z} \right)} = {\frac{\mu\; R^{2}}{\pi\;\rho}{\sin(\varphi)}{\int_{0}^{\infty}{k^{2}{\mathbb{d}{k\left( {{{\theta\left( {\rho - R} \right)}{I_{n}^{\prime}({kR})}{K_{n}\left( {k\;\rho} \right)}} + {{\theta\left( {R - \rho} \right)}{K_{n}^{\prime}({kR})}{I_{n}\left( {k\;\rho} \right)}}} \right)}}{F_{C}\left( {k,z} \right)}}}}}} & \; \\{\mspace{79mu}{{n = 0},1,2}} & \;\end{matrix}$

The Bx-component of the magnetic flux field inside the cylinder of thecoil 120′ is:

${B_{X}\left( {\rho,\varphi,z} \right)} = {{{- \frac{\mu\; R^{2}}{2\pi}}{\int_{0}^{\infty}{k^{2}{\mathbb{d}{{kI}_{0}\left( {k\;\rho} \right)}}{K_{1}^{\prime}({kR})}{F_{C}\left( {k,z} \right)}}}} - {\frac{\mu\; R^{2}}{2\;\pi}{\cos\left( {2\;\varphi} \right)}{\int_{0}^{\infty}{k^{2}{\mathbb{d}{{kI}_{2}\left( {k\;\rho} \right)}}{K_{1}^{\prime}({kR})}{F_{C}\left( {k,z} \right)}}}}}$

This Bx-component (in the local system of coordinates) should cancel theBz'-component produced by the magnet. This suggests that a minimizationprocedure can be applied (similar to that of the gradient design) tofind the currents density f_(Z)(z). We consider a functional to beminimized:

$W = {E + {\frac{\Lambda}{2\;\mu}{\sum\limits_{i \in V_{Linac}}^{\;}\;\left\lbrack {{B_{X}^{Coil}\left( r_{i} \right)} - B_{Z,i}^{Magnet}} \right\rbrack^{2}}} + {\frac{\beta}{2\;\mu}{\sum\limits_{i \in {DSV}}^{\;}\;\left\lbrack {B_{X}^{Coil}\left( r_{i} \right)} \right\rbrack^{2}}} + {\frac{\mu\;\lambda}{2}{\frac{\partial^{k}{f_{Z}(z)}}{\partial z^{k}}}_{2}^{2}}}$

In the above equation, E is the energy of the coil 120′, the second termis to minimize the deviation of the field produced by the shielding coil120′ from that of the main magnets 112 a, 112 b, the third term is tominimize the effect of the shield coil 120′ on the field in-homogeneityin the imaging volume, and the last term is introduced as to limit thecurrent density. The coefficients Λ, β, and λ are the weighting factors;λ can be a regularization parameter to minimize the current in theshielding coil 120′.

The current density f_(Z)(z) can be expressed in terms of a basisfunctions. It should be mentioned that the current density f_(Z)(z) iszero at the ends of the shielding coil 120′.

${f_{Z}(z)} = {{\sum\limits_{n = 1}^{\;}\;{a_{n}{\Phi_{n}^{C}\left( {z,L} \right)}}} + {\sum\limits_{n = 1}^{\;}\;{b_{n}{\Phi_{n}^{S}\left( {z,L} \right)}}}}$Φ_(n)^(C)(z, L) = cos (k_(n)^(C)z), Φ_(n)^(S)(z, L) = sin (k_(n)^(S)z)${k_{n}^{C} = \frac{\pi\left( {{2\; n} - 1} \right)}{2\; L}},{k_{n}^{S} = \frac{\pi\; n}{L}}$

The coefficients α_(n) can be found from the following equation:

${\frac{\partial W}{\partial a_{n}} = 0},{\frac{\partial W}{\partial b_{n}} = 0}$

This is leading to a system of linear equation for the coefficientsα_(n). The energy E has the following form:

$E = {{{- \frac{\mu\; R^{2}}{2}}{\int_{0}^{\infty}\ {{\mathbb{d}{k({kR})}^{2}}{I_{1}^{\prime}({kR})}{K_{1}^{\prime}({kR})}\text{<}f_{Z}{\cos\left( {k{()}} \right)}f_{Z}\text{>}}}} = {\frac{\mu}{2}{\sum\limits_{n,{m = 1}}^{\;}\;{A_{n}W_{n,m}A_{m}}}}}$$\mspace{20mu}{A_{\alpha} = \begin{pmatrix}a_{n} \\b_{m}\end{pmatrix}}$     W_(α, β) = −R²∫₀^(∞) 𝕕k(kR)²I₁^(′)(kR)K₁^(′)(kR)<Φ_(α)cos (k())Φ_(β)>$\mspace{20mu}{\Phi_{\alpha} = \begin{pmatrix}\Phi_{n}^{S} \\\Phi_{n}^{C}\end{pmatrix}}$     <Φ_(α)cos (k())Φ_(β)> = ∫_(L)^(L₂)∫_(L)^(L₂)Φ_(α)(z)cos (k(z − z^(′)))Φ_(β)(z^(′)) 𝕕z 𝕕z^(′)

The field produced by the shield coil 120′ has the following form:

$\mspace{79mu}{{B_{X}^{Coil}(r)} = {\mu{\sum\limits_{\alpha}^{\;}\;{A_{\alpha}{B_{X,\alpha}^{Coil}(r)}}}}}$${B_{X,\alpha}^{Coil}(r)} = {{{- \frac{R^{2}}{2\;\pi}}{\int_{0}^{\infty}{k^{2}\ {\mathbb{d}{{kI}_{0}\left( {k\;\rho} \right)}}{K_{1}^{\prime}({kR})}{\Psi_{C,\alpha}\left( {k,z} \right)}}}} - {\frac{R^{2}}{2\;\pi}{\cos\left( {2\;\varphi} \right)}{\int_{0}^{\infty}{k^{2}\ {\mathbb{d}{{kI}_{2}\left( {k\;\rho} \right)}}{K_{1}^{\prime}({kR})}{\Psi_{C,\alpha}\left( {k,z} \right)}}}}}$  Ψ_(C, α)(k, z) = ∫_(L₁)^(L₂)Φ_(α)(z^(′))cos (k(z − z^(′))) 𝕕z^(′)$\mspace{20mu}{{\frac{\partial^{k}{f_{Z}(z)}}{\partial z^{k}}}_{2}^{2} = {{L{\sum\limits_{n = 1}^{\;}\;{a_{n}^{2}\left( k_{n}^{C} \right)}^{2\; k}}} + {b_{n}^{2}\left( k_{n}^{S} \right)}^{2\; k}}}$

Then the equation for the unknown A holds:

$\mspace{79mu}{{\sum\limits_{\beta}^{\;}\;{Z_{\alpha,\beta}A_{\beta}}} = K_{\alpha}}$$Z_{\alpha,\beta} = {W_{\alpha,\beta} + {\alpha{\sum\limits_{i \in V}^{\;}\;{{B_{X,\alpha}^{Coil}\left( r_{i} \right)}{B_{X,\beta}^{Coil}\left( r_{i} \right)}}}} + {\beta{\sum\limits_{i \in {DSV}}^{\;}\;{{B_{X,n}^{Coil}\left( r_{i} \right)}{B_{X,m}^{Coil}\left( r_{i} \right)}}}} + {\lambda\;\delta_{\alpha,\beta}\frac{k_{n}^{2\; k}}{2}L}}$$\mspace{20mu}{K_{\alpha} = {\frac{\Lambda}{\mu}{\sum\limits_{i \in V}^{\;}{{B_{X,\alpha}^{Coil}\left( r_{i} \right)}{B_{Z}^{Magnet}\left( r_{i} \right)}}}}}$

The matrix Z_(α,β) is positive defined and does not have zeroeigenvalue, thus:

$a_{\alpha} = {\sum\limits_{\beta}^{\;}\;{\left( Z^{- 1} \right)_{\alpha,\beta}K_{\beta}}}$

This defines the solution for the current density.

Some embodiments can include a combined passive shield and active coil.The residual Bz-field shown in FIG. 5b (a single shell case) was used asan input data. The radius of the coil 120′ was chosen to be 75 mm andthe half length L was chosen to be L=180 mm. The center of the coil 120′is located at y=1051 mm. FIG. 12A shows the z-component of the currentdensity on the active shield coil 120′ prior to activation (i.e., priorto application of an electric current) of the coil 120′, and FIG. 12Bshows the residual Bz-field after activation of the shield coil 120′.

The following parameters Λ, γ, K, and λ were used: κ=1, ρ=0, K=1, andλ=0.0001. The parameter β that accounts for the effect of correcting thein-homogeneity inside the DSV was chosen to be zero because the level ofthe residual field of FIG. 5B is already small (of the order of 7 Gauss)and the active shield coil is located far from the imaging volume.

Some embodiments can include a completely active coil shielding system.In such embodiments, the shielding of the linac 107 can be accomplishedlocally using only the above-described active current-carrying coils,such as coil 120′, in place of the passive magnetic shields inembodiments described above. The coils 120′ can be arranged to simplycancel the field at the linac 107 and can also incorporate an activeshield to reduce the influence on the homogeneity of the main magneticfield.

Still another alternative way of shielding the linac 107 locally is touse a distribution of permanent magnets. They can be arranged to simplycancel the field at the linac 107 and can also incorporate an activeshield, such as coil 120′, to reduce the influence on the homogeneity ofthe main magnetic field from the main magnets 112 a and 112 b.

All possible combinations of the disclosed embodiments are alsopossible. Small variations to the shields and distributions of theshielding materials, current carrying coils, and magnet distributionsare also possible.

It should be noted that the magnetic shields described herein, such asshields 118, 120, 122, 130, 132, and others experience a force from themain magnets 112 a and 112 b of the MRI 102. Thus, the mounting for theshields is preferably designed to withstand such magnetic forces.

The high-power RF source and waveguide for the linac 107 can also beenclosed, or partially enclosed, within the magnetic shields disclosedherein. The RF shielding can be extended to contain some or allcomponents of the linac 107.

Regarding RF screening for the MRI 102, clinical linacs suitable for useas linac 107 can operate in the S-band frequency range accelerateelectrons to about 6 MeV using RF microwave cavities at ˜3 GHz. Whilethis frequency is well above the 15 MHz of the MRI system 102, itinvolves megawatts of RF power pulse with a frequency of several hundredHertz. Sidebands in the RF power source can excite/reflect from othermaterials causing interference with the operation of the MRI system 102.As mentioned above in connection with FIG. 4B, the element 120 can be anRF shield that is placed around the linac 107 made of RF absorbing, RFreflecting, or a combination of both can effectively eliminate the RFinterference with the MRI system 102. Additionally, the MRI RF room,which can be made of RF reflecting material that can bound RF from thelinac 107 into the MRI 102, can be lined on the interior surface with awall covering of RF absorbing material, e.g., meshed or chopped carbonfiber, carbon fiber wallpaper, carbon fiber panels, or carbon fiberpaint, and eliminate RF reaching the MRI. The gantry 106 and area aroundthe RF source of the linac 107 can be covered in RF absorbers,reflectors, and combinations of both to reduce the ambient(environmental) RF fields. At 3 GHz (microwave ovens are at 2.45 GHz)the RF will produce dielectric heating of polarized molecules such aswater. Thus, a variety of polarized molecule materials can be used as RFabsorber for the RF energy. In a split magnet system, some of theconductive surfaces that divert RF energy in a closed system are missingin the magnet gap 114. An RF shield about the MRI bore can be used inconjunction with the other shielding method described above. The RFshields do not add significantly to the beam attenuation so that thequality of the radiotherapy is significantly compromised. The conductiveshielding may or may not be grounded to the magnet. If these surfaceswere made of aluminum, such as aluminum foil, the beam attenuation wouldeven be less than using copper. If the gradient coil is wound on aformer one can construct the former out of carbon fiber for isolationfrom the linac system.

To aid in reading the drawings, the following listing of element numbersand their respective components is provided:

-   -   100 Radiation therapy system    -   102 MRI device    -   104 Radiation source    -   106 Gantry    -   107 Linear accelerator    -   108 Patient couch    -   107 a Waveguide    -   107 b Electron gun    -   107 c Target    -   107 d Pre-collimator    -   107 e Multi-leaf collimator    -   107′ Second linear accelerator    -   110 Patient    -   112 a Main magnet    -   112 b Main magnet    -   114 Central gap    -   115 Cooling system    -   118 Outer magnetic shield    -   118 a Bottom edge    -   120 Inner magnetic shield    -   120′ Shielding coil    -   EB Electron beam    -   CP Central axial plane    -   IC Isocenter    -   P1 Point    -   P2 Point    -   RD Rotation directions    -   120A Slot    -   120B Slot    -   122 Magnetic shield device    -   126 Split radiotherapy magnet    -   128 Split radiotherapy magnet    -   130 Isolated shell    -   132 Isolated shell    -   140 Second shield    -   144 Annulus disc    -   146 Annulus disc

While various embodiments in accordance with the disclosed principleshave been described above, it should be understood that they have beenpresented by way of example only, and are not limiting. Thus, thebreadth and scope of the invention(s) should not be limited by any ofthe above-described exemplary embodiments, but should be defined only inaccordance with the claims and their equivalents issuing from thisdisclosure. Furthermore, the above advantages and features are providedin described embodiments, but shall not limit the application of suchissued claims to processes and structures accomplishing any or all ofthe above advantages.

Additionally, the section headings herein are provided for consistencywith the suggestions under 37 C.F.R. 1.77 or otherwise to provideorganizational cues. These headings shall not limit or characterize theinvention(s) set out in any claims that may issue from this disclosure.Specifically and by way of example, although the headings refer to a“Technical Field,” such claims should not be limited by the languagechosen under this heading to describe the so-called technical field.Further, a description of a technology in the “Background” is not to beconstrued as an admission that technology is prior art to anyinvention(s) in this disclosure. Neither is the “Summary” to beconsidered as a characterization of the invention(s) set forth in issuedclaims. Furthermore, any reference in this disclosure to “invention” inthe singular should not be used to argue that there is only a singlepoint of novelty in this disclosure. Multiple inventions may be setforth according to the limitations of the multiple claims issuing fromthis disclosure, and such claims accordingly define the invention(s),and their equivalents, that are protected thereby. In all instances, thescope of such claims shall be considered on their own merits in light ofthis disclosure, but should not be constrained by the headings set forthherein.

What is claimed is:
 1. A shielded linear particle accelerator (linac)system comprising: one or more magnets positioned along a longitudinalaxis and configured to generate a magnetic field, at least a portion ofthe magnetic field extending along the longitudinal axis; a linearparticle accelerator configured to direct a radiation beam towards thelongitudinal axis; and a shield configured to shield the linac from themagnetic field, the shield comprising at least a first shell and asecond shell, the first shell and the second shell at least partiallysurrounding the linac, the second shell separate from and nested withinthe first shell.
 2. The shielded linac system of claim 1, wherein thefirst shell at least partially surrounds the linac and the nested secondshell.
 3. The shielded linac system of claim 1, wherein the first shelland the second shell are magnetically and electrically isolated fromeach other.
 4. The shielded linac system of claim 1, wherein the firstshell and the second shell are magnetically isolated from each other. 5.The shielded linac system of claim 1, wherein the first shell and thesecond shell are formed of any combination of steel, copper, aluminum,and carbon fiber.
 6. The shielded linac system of claim 1, wherein thefirst shell and the second shell are concentric about a common axis. 7.The shielded linac system of claim 1, wherein the first shell is formedof a first material that includes at least one of an RF absorbingmaterial and an RF reflecting material, and wherein the second shell isformed of a second material that includes at least one of the RFabsorbing material and the RF reflecting material.
 8. The shielded linacsystem of claim 1, wherein the one or more magnets comprise a firstmagnet and a second magnet, and wherein the longitudinal axis extendsthrough an isocenter of the first magnet and the second magnet.
 9. Theshielded linac system of claim 8, wherein the linac is positioned at afixed radial distance from the isocenter of the first magnet and thesecond magnet.
 10. The shielded linac system of claim 1, furthercomprising: a gantry rotatable about the longitudinal axis, wherein thelinac is supported by the gantry and configured to rotate with thegantry.
 11. A linear particle accelerator (linac) shielding methodcomprising: generating, with a magnetic resonance imaging (MRI) system,a magnetic field by one or more magnets positioned along a longitudinalaxis of the MRI system, at least a portion of the magnetic fieldextending along the longitudinal axis of the MRI system; and shielding alinear particle accelerator (linac) from the magnetic field using ashield, wherein the shield comprises at least a first shell and a secondshell, the first shell and the second shell at least partiallysurrounding the linac, the second shell separate from and nested withinthe first shell.
 12. The linac shielding method of claim 11, wherein thefirst shell at least partially surrounds the linac and the nested secondshell.
 13. The linac shielding method of claim 11, wherein the firstshell and the second shell are magnetically and electrically isolatedfrom each other.
 14. The linac shielding method of claim 11, wherein thefirst shell and the second shell are magnetically isolated from eachother.
 15. The linac shielding method of claim 11, wherein the firstshell and the second shell are formed of any combination of steel,copper, aluminum, and carbon fiber.
 16. The linac shielding method ofclaim 11, wherein the first shell and the second shell are concentricabout a common axis.
 17. The linac shielding method of claim 11, whereinthe first shell is formed of a first material that includes at least oneof an RF absorbing material and an RF reflecting material, and whereinthe second shell is formed of a second material that includes at leastone of the RF absorbing material and the RF reflecting material.
 18. Thelinac shielding method of claim 11, wherein the one or more magnetscomprise a first magnet and a second magnet, and wherein thelongitudinal axis extends through an isocenter of the first magnet andthe second magnet.
 19. The linac shielding method of claim 18, whereinthe linac is positioned at a fixed radial distance from the isocenter ofthe first magnet and the second magnet.
 20. The linac shielding methodof claim 11, wherein the linac is supported by a gantry and configuredto rotate with the gantry and direct a radiation beam towards thelongitudinal axis.